The present invention relates to image-enhancing agents, contrast agents or spectral shift agents to enhance tissue or organ images or nuclear spectra obtained from live animals with magnetic resonance imaging (MRI) or spectroscopy (MRS), radioisotope scanning or ultrasound imaging.
Magnetic resonance imaging is a medical procedure that takes advantage of the magnetic spin properties of nuclei to create an image fundamentally like that of the more widely known X-ray procedure the CT scan. The nucleus of an atom contains neutrons and protons. Atoms which have an odd number of neutrons or protons have a non-zero spin quantum number (I). These atoms can be thought to behave like small spinning spheres. Because the nuclei have positive charges, the spinning produces a magnetic moment (U), analogous to an electric current in a closed loop of wire. When exposed to a magnetic field these dipoles align themselves with the magnetic field. The spin (magnetic moment) vectors experience a torque when subjected to a magnetic field. Due to this torque, the nuclei process about the axis of the magnetic field at a rate given by the Larmor relationship: EQU f=w/2.pi.=.gamma.Bo/2.pi.
f=resonance frequency in Hertz (Hz) PA1 w=the angular frequency in radians per second PA1 .gamma.=magnetic gyric ratio PA1 Bo=static magnetic field PA1 .gamma.=a nuclear constant characteristic of the isotope PA1 B.sub.1 =amplitude of the FR field PA1 .pi.=duration PA1 N(H)=number of protons in the discrete tissue volume (spin density); PA1 f(v)=a function of proton velocity and the fraction of protons which are moving (e.g., due to following blood); PA1 TE=time between the radio frequency (rf) pulse and the detection of signal (spin-echo); PA1 TR=the interval between repetition of the rf pulse; PA1 T1=the time interval associated with the rate of proton energy transfer to the surrounding chemical environment (spin-lattice relaxation); PA1 T2=the time interval associated with the rate of proton energy transfer, one to other (spin-spin relaxation).
The resonance frequency of a nucleus is a function of its magnetogyric ratio (a constant for the particular nucleus being studied) and the strength of an applied magnetic field (Larmor equation). The magnetogyric ratio of a nucleus relates the magnetic moment and the nuclear spin quantum number (I). The spin quantum number depicts the number of energy states that a nucleus can have; a nucleus has 2I+1 energy levels. The hydrogen nucleus (a proton) has a nuclear spin of 1/2, and thus two possible spin states. In nuclear magnetic resonance spectroscopy (NMR), the magnetic moment is shown as revolving around a fixed magnetic field (FIG. 1) at a fixed frequency. In a sample, large numbers of nuclei are revolving around this field in one of the two possible spin states, creating a "macroscopic magnetization" M parallel to the magnetic field. There are more nuclei in the low-energy spin state when there is no outside influence (FIG. 2).
Resonance is the induction of a transition between two different energy states. The nuclei that lends itself best to magnetic resonance imaging is the proton, the major isotope of hydrogen. Hydrogen has a very large abundance in biological systems in water (E10.sup.23 /cm.sup.3) as well as in other biochemical molecules and has the required magnetic moment. The energy necessary to produce a transition between the two spin states of hydrogen (+1/2 and -1/2) is the difference in energy (.DELTA.E) between these spin states. In MRI, resonance occurs when radiofrequency (RF) energy is applied at the Larmor frequency, flipping the magnetic moments from their m=+1/2 (lower energy) to their m=-1/2 (higher energy) states. Magnetic resonance absorption can only be detected by transverse magnetization (magnetization perpendicular to Bo). Only the transverse component, Mxy, is time dependent and therefore according to Faraday's law of induction, only time dependent magnetization can induce a voltage in a receiver coil. Transverse magnetization is generated when a radiofrequency (RF) field of amplitude B., rotating synchronously with the processing spins is applied.
When the RF field acts in a direction perpendicular to the main field, the effect is to rotate the magnetization away from the rest state. The macroscopic magnetization experiences a torque of the RF field, forcing the magnetization to rotate about it. If the duration of the B. field is such that the net magnetization is rotated by an angle of 90.degree., it will become transverse or perpendicular to the stated field (FIG. 3). The angle of rotation .theta., the RF flip angle, is given by the Ernst equation. EQU .theta.=.gamma.B.sub.1 .tau.
Once the RF field is removed, the magnetization is subjected to the effect of the static magnetic field and processes about it. With a detection coil positioned with its axis along the y axis, the AC voltage induced in the coil is given by EQU .gamma..delta.Mxy.degree.cos w.tau.
Mxy.degree.=initial transverse magnetization following a 90 degree RF pulse PA0 .tau.=time interval between the rotation PA0 1) inherent spin density; PA0 2) longitudinal relaxation time (T1); PA0 3) transverse relaxation time (T2); and PA0 4) flow. PA0 1) the paramagnetic concentration; PA0 2) the distance (.gamma..sup.-6); and PA0 3) a time constant describing the dynamic nature of electron-proton interactions (correlation time). PA0 1) increased effective magnetic moment; PA0 2) decreased freedom of molecular motion; and PA0 3) decreased water .sup.1 H exchange. PA0 1. They increase the specificity of MRI diagnosis. PA0 2. Smaller lesions can be identified earlier. PA0 3. Image-enhancing agents enhance tumor masses differently than surrounding edema fluid or abscesses. This allows the extent and invasion of tumor to be defined more precisely. Lesion with infiltrative-type growth (e.g., certain metastatic carcinomas and glioblastomas) will require contrast agents for demarcation between tumor and edema fluid (Felix et al. (1985) Proc. Soc. Mag. Res. Med. V 2, p 831). PA0 4. Image-enhancing agents improve the distinction between recurrent tumor and fibrous tissue resulting from surgery and radiation. PA0 5. Image-enhancing agents can decrease the time required per scan and potentially decrease the number of scans required per procedure. This increases the volume of procedures and decreases their expense. PA0 6. Body imaging has a significantly lower resolution (typically 0.5-1.0 cm) and sensitivity (decreased signal-to-noise ratio) than brain imaging (Wesbey et al. (1983) Radiology V 149, p 175). These differences result from the greater inhomogeneity of the magnetic field; the larger radio frequency coil; unequal phase-pulsing of deep versus shallow nuclei; and motion artifacts produced by respiration, cardiac systole, gastrointestinal peristalsis, and voluntary muscle movement; and PA0 7. Advanced (polymeric and microsphere) forms of contrast agents (see below) appear to be required for the optimal acquisition and interpretation of blood-flow and tissue-perfusion images and related spectral (phase) information.
The transverse magnetization decays to zero exponentially with a time constant T2*.
Therefore: EQU .gamma..delta.Mxy.degree.e.sup.-.tau./T2 * cos w.tau.
This equation represents a damped oscillation that is called the induction decay or F10 signal. As the transverse magnetization decays to zero with time, the longitudinal magnetization increases back to its equilibrium value. This return to equilibrium values is termed relaxation. RF stimulation causes the nuclei to absorb energy transferring them to the excited state. The nuclei can return to the ground state by transferring energy to their surroundings, the so called lattice. This method of relaxation is called spin-lattice relaxation (T1). (see FIG. 4) Components of M, after being rotated to the x'y' plane, (FIG. 5) return to their original magnetization values in a time (T2), the "spin-spin" relaxation time. In the T2 relaxation process, nuclei in the excited and ground state exchange energy with each other. Thus T2 measures the amount of time necessary for the nuclei to get out of phase with each other and return to the original random state. Both T1 and T2 relaxation times of a nuclei can vary widely from milliseconds to minutes depending upon the nuclei and the environment surrounding them. The primary task of magnetic resonance imaging is soft tissue contrast and the detection of low-contrast lesions. The detection of lesions depends on the inherent difference in contrast between lesions and surrounding normal tissues. The tissue or media present around a resonating nucleus have differential effects that can alter the T1 and T2 relaxation times. However, in general, organic substances such as those found in the body (tissues, organs) have a fairly uniform effect on relaxation times, largely due to the high percentage of water. MRI scans of a substance take advantage of relaxation times (T1 and T2) differences to generate an image of the object being scanned in "slices." The difference in relaxation times of, say, different organs in the body allows a visible image to be formed. However, the largely uniform effect on relaxation times of most parts of the body causes an image that is very difficult to see due to lack of contrast, hence the need for "contrast agents."
Inherent tissue contrast in MRI is determined by differences in:
Contrast agents can improve visualization of low contrast organs and lesions. The most promising agents affect signal by enhancing relaxation. A contrast agent is a substance, either 1) a paramagnetic metal ion, 2) free oxygen, or 3) a substance with free radicals (unpaired electrons), that has a far different effect on proton relaxivity than water. A good contrast agent can be either directly injected into the target area or tagged to an antibody against the target area (e.g., a cancerous tumor) and thus provide sharp contrast for aid in viewing the area by MRI.
Pharmacologic basis for relaxation enhancement is based on positive magnetic susceptibility (Bourdreaux, E. A. and Mulap, L. N.: THEORY AND APPLICATION OF MOLECULAR PARAMAGNETISM, New York, 1976, John Wiley and Sons). When a substance is placed in an external magnetic field, induced magnetization in the substance is additive to that of the applied field. The magnetic susceptibility of a substance is defined as the ratio of induced magnetization to that of the applied field. Substances can be categorized by their magnetic susceptibility (see Table I).
TABLE I ______________________________________ (David Stark and William Bradley: MAGNETIC RESONANCE IMAGING, St. Louis, 1988, Mosby) CLASS BASIS SUSCEPTIBILITY ______________________________________ Diamagnetic paired electrons- -10.sup.-6 no permanent spin moment Paramagnetic unpaired electrons +10.sup.-1 non-interacting permanent moments Superparamagnetic unpaired electrons- +10.sup.+2 non-interacting domains Ferromagnetic unpaired electrons- +10.sup.+2 interacting domains ______________________________________
Diamagnetic substances have negative susceptibilities. Most organic and inorganic compounds are diamagnetic and since all atoms experience induced magnetic affects arising from electron orbital motion, a diamagnetic component is present in all materials. Diamagnetic effects are very weak and can be overwhelmed in magnitude by relatively few unpaired electron spins. Diamagnetic materials are generally of little interest as contrast agents.
Paramagnetic, superparamagnetic, and ferromagnetic substances are characterized by the predominant magnetic effects of unpaired electron spins which produce positive susceptibilities and positive induced magnetization.
Paramagnetism is characterized by independent action of individual atomic or molecular magnetic moments. Ferromagnetism is characterized by solid phase microscopic volumes or domains in which unpaired electron spins are permanently aligned. Multiple domains in bulk can be isotropic (unmagnetized) or anisotropic (magnetized).
Superparamagnetic materials can be regarded as single domain particles. The susceptibilities per atom or mole of these substances exceed those of corresponding soluble paramagnetic species due to magnetic ordering.
Superparamagnetic and ferromagnetic susceptibilities increase linearly with field strength. Superparamagnetic substances unlike ferromagnetic substances are characterized by restoration of induced magnetization to zero upon removal of the external field.
Paramagnetic enhancement of nuclear relaxation was first described in 1946 by Bloch and co-workers (Bloch, F., Hansen, W. N., and Packard, P.: "The nuclear induction experiment", Physiol. Rev. 70:474-485, 1946) when they demonstrated a convenient practice of shortening the time needed to observe water .sup.1 H T1 by adding ferric nitrate (a paramagnetic solute).
Positive susceptibility is necessary, but not sufficient, for effective relaxation. The magnitude of relaxation enhancement also depends on proximity and on correlation time. A mathematical formulation of paramagnetic enhanced solvent relaxation is described by Bloembergen (Bloembergen, N.: "Proton relaxation times in paramagnetic solutions", J. Chem. Phys. 27:572, 1957). These equations imply that nuclear relaxation results from several simultaneous mechanisms.
The paramagnetic contribution to nuclear relaxation is proportional to
The correlation time is dominated by the fastest rate of paramagnetic tumbling, electron spin flips, or chemical exchange. Due to more optimal correlation of spin motion, nuclear T1 relaxation enhancement in biologic systems is more effective with relaxation agents if large molecular weight or asymmetric shape, that is, relatively long rotational correlation times.
Solvent relaxation is the presence of superparamagnetic particles chiefly differs from that in the presence of paramagnetic solutes due to much greater weighting of the magnetic moment contribution. Compared with paramagnetic solutes, superparamagnetic particulates have
The much greater effective magnetic moment dominates these factors and results in T2 shortening caused by long range effect from magnetic field heterogeneity.
The imaging of internal structures and organs of live animals has been an important aspect of medicine since the advent of X-ray usage for this purpose. Among the techniques more recently developed for such imaging are those involving scanning for emission of particles form an internally located radioisotope. Such radioisotopes preferably emit gamma particles and are generally isotopes of metallic elements. One problem common to the diagnostic usage of such gamma particle-emitting radioisotopes concerns the localization of these materials at sites of particular interest rather than to have them randomly dispersed or rapidly excreted, by the kidney, for example. Another problem of such radioisotope mediated imaging concerns optimizing the circulating half-life of radioisotopes, for example, by preventing or accentuating their binding to serum proteins (e.g., albumin), or by prior conjugation (complexation) to polymeric carriers or receptor-binding substances.
A second class of internal body imaging which is undergoing a rapid growth in clinical use is ultrasound imaging. This is based on the detection of differences in the internal velocity (reflectivity) of directed, high-frequency sound waves. Differences in image brightness are produced at the interfaces between tissues with different native densities and ultrasound reflectivities. A present clinical problem is the difficulty of visualizing lesions in the stomach, small and large bowel, bladder, and cavities of the female reproductive tract, due to similarities of ultrasound velocity between these organs of interest and immediately adjacent tissues. Diagnostic introduction of a dense, nonradioactive metal element or ion at sufficient concentrations can confer the significant differences in ultrasound reflectivity which are required to visualize otherwise undetectable tumors and inflammatory lesions.
NMR intensity and relaxation images have been shown in recent years to provide a third important method of imaging internal structures and organs of live animals. Clinical magnetic resonance Imaging (MRI) is a rapidly growing, new form of brain and body imaging. Low-field (proton) MRI detects chemical parameters in the immediate environment around the protons because of body tissues (predominantly water protons because of their relative abundance). Changes in these parameters occur very early in disease and are independent of physical densities detected by ionizing radiation. In the brain and central nervous system, MRI has allowed detection of tumors at an earlier clinical stage and with fewer imaging artifacts than is possible with computerized axial tomography (CAT) (Runge et al., (1983) Am. J. Radiol. V 141, p 1209). Under optimal conditions, image resolution is in the submillimeter size range.
Seven factors are among those making it important to develop nontoxic MRI image-enhancing agents analogous to those available for CAT.
The discrete intensities of a two-dimensional, Fourier-transformed image are described by the following general equation (for spin-echo pulse sequences): EQU Intensity=N(H).multidot.f(v).multidot.exp(-TE/T2).multidot.(1-exp(TE-TR)/T1 ),
where
The T1 and T2 times have reciprocal effects on image intensity. Intensity is increased by either shortening the T1 or lengthening the T2 or vice versa depending upon whether proton density, T1-weighted or T2-weighted images are desired. Tissue contrast occurs naturally and is related to variations in the chemical environments around water protons (major contributor) and lipid protons (usually minor). Chemical agents have been used to enhance this natural contrast. The one most widely tested clinically is the paramagnetic metal ion, gadolinium (Gd.sup.+3) chelated to an appropriate organic chelate (Runge et al. (1983) Am. J. Radiol. V 142, p 619). Although gadolinium shortens both the T1 and T2 times, at the lower does used for clinical imaging, the T1 effect generally predominates and the image becomes brighter. Also, the rf pulse sequence can be programmed to accentuate T1 changes and diminish those due to T2 (Runge et al. (9183) Am. J. Radiol. V 141, p 1209). Hence, "T1-weighted" enhancement can be achieved by selecting the most favorable Gd dose and rf pulse sequence.
The shortening of proton relaxation times by Gd is mediated by dipole-dipole interactions between its unpaired electrons and adjacent water protons. The effectiveness of Gd's magnetic dipole drops off very rapidly as a function of its distance form these protons (as the sixth power of the radius) (Brown (1985) Mag. Res. Imag. V 3, p 3). Consequently, the only protons which are relaxed efficiently are those able to enter Gd's first or second coordination spheres during the interval between the rf pulse an signal detection. This ranges from 10.sup.5 to 10.sup.6 protons/second (Brown (1985) Mag. Res. Imag. V 3, p 3). Still, because Gd has the largest number of unpaired electrons (seven) in its 4f orbital, it has the largest paramagnetic dipole (7.9 Bohr magnetons) and exhibits the greatest paramagnetic relaxivity of any element (Runge et al. (1983) Am. J. Radiol. V 141, p 1209 and Weinman et al. (1984) Am. J. Radiol. V 142, p 619). Hence, Gd has the highest potential of any element for enhancing images. However, the free form of Gd is quite toxic. This results in part from precipitation at body pH (as the hydroxide). In order to increase solubility and decrease toxicity, Gd has been chemically chelated by small organic molecules. To date, the chelator most satisfactory from the standpoints of general utility, activity, and toxicity is diethylenetriamine pentaacetic acid (DTPA) (Runge et al. (1983) Am. J. Retail V 141, p 1209 and Weinman et al. (1984) Am. J. Retail V 142, p 619). The first formulation of this chelate to undergo extensive clinical testing was developed by Schering Ag - Berlex Imaging according to a patent application filed in West Germany by Gries, Rosenberg and Weinmann (DE-OS 3129906 A 1 (1981). It consists of Gd-DTPA which is neutralized and stabilized with the organic base, N-methyl-D-glucamine (meglumine). The Schering-Berlex agent is nearing completion of Phase III clinical testing at selected centers across the United States and abroad. The results of preliminary studies indicated that almost all human brain tumors undergo significant enhancement (Felix et al. (1985) Proc. Soc. Mag. Res. Med. V 2, p 831 and K. Maravilla, personal communication). These include metastatic carcinomas, meningiomas, gliomas, adenomas and neuromas. Renal tumors are also enhanced satisfactorily (Lanaido et al. (1985) Proc. Soc. Mag. Res. Med. V 2, p 877 and Brasch et al. (183) Am. J. Retail. V 141. p 1019). The Schering-Berlex formulation was available for general clinical use by 1989.
Despite its satisfactory relaxivity and toxicity, this formulation has four major disadvantages.
(1) Chelation of Gd markedly decreases its relaxivity (by 1/2 an order of magnitude). This happens because chelators occupy almost all of Gd's inner coordination sites which coincide with the strongest portion of the paramagnetic dipole (Koenig 1985) Proc. Soc. Mag. Res. Med. V 2, p 833 and Geraldes et al. (1985) Proc. Soc. Mag. Res. med. V 2, p 860).
(2) Gd-DTPA dimeglumine, like all small paramagnetic metal chelates, suffers a marked decrease in relaxivity at the higher radio frequencies used clinically for proton imaging (typically 5 MHz, 2T) (Geraldes et al. (1985) Proc. Soc. Mag. Res. med. V 2, p 860).
(3) Due to its low molecular weight, Gd-DTPA dimeglumine is cleared very rapidly from the bloodstream (t.sup.1/2 in 20 minutes) and also from tissue lesions (tumors) (Weinman et al. (1984) Am. J. Radiol V 142, p 619). This limits the imaging window (to ca. 30 to 45 minutes); limits the number of optimal images after each injection (to ca. 2); and increases the agent's required dose and relative toxicity.
(4) The biodistribution of Gd-DPTA is suboptimal for imaging of body (versus brain) tumors and infections. This is due to its small molecular size. Intravenously administered Gd-DTPA exchanges rapidly into the extracellular water of normal tissues, as well as concentrates in tumors and infections. This is facilitated by an absence in body organs, of the "blood-brain" vascular barrier which partly restricts the exchange of Gd-DTPA into the extracellular water of normal (versus diseased) brain. The result in body organs, is a reduced difference in the concentration of Gd-DTPA between normal and diseased regions of tissue, and hence, reduced image contrast between the normal and diseased regions of the organ. Also a disproportionate quantity (&gt;90%) of Gd-DTPA is sequestered very rapidly in the kidneys (Weinman et al. (1984) Am. J. Radiol V 142, p 619). Of much greater interest to body MRI, are the abdominal sites involved in the early detection and staging of tumors (particularly the liver, and also the spleen, bone marrow, colon and pancreas).
Three approaches have been taken in attempts to overcome these disadvantages.
(1) Alternative, small chelating molecules have been tested. These make Gd more accessible to water protons but still chelate the metal with a sufficient affinity to potentially control its toxicity in vivo. The most effective of these chelators is DOTA, the poly-azamacrocyclic ligand, 1,4,7,10-tetraazacyclododecane-N,N',N"-tetraacetic acid (Geraldes et al. (1985) Proc. Soc. Mag. Res. Med. V 2, p 860). Its relaxivity is approximately 2 times greater than that of Gd-DTPA over a wide range of Larmor frequencies. However, it is still less active than free Gd.
(2) Gd and Gd-chelates have been chemically conjugated to macromolecules, primarily the proteins, albumin (Bulman et al. (1981) Health Physics V 40, p 228 and Lauffer et al. (1985) Mag. Res. Imaging V 3, p 11), asialofetuin (Bulman et al. (1981) Health Physics V 40, p 228), and immunoglobulins (Lauffer et al. (1985) Mag. Res. Imaging V 3, p 11 and Brady et al. (1983) Soc. Mag. Res., 2nd Ann. Mtg., Works in Progress, San Francisco , Calif.). This increases the relaxivity of Gd by slowing its rate of molecular tumbling (rotational correlation time) (Lauffer et al. (1985) Mag. Res. Imaging V 3, p 11). This improves coupling of the energy-transfer process between protons and Gd (Geraldes et aI. (1985) Proc. Soc. Mag. Res. Med. V 2, p 860, Lauffer et al. 91985) Mag. Res. Imaging V 3, p 11 and Brown et al. (1977) Biochemistry V 16, p 3883). Relaxivities are increased by multiples of 5 to 10 relative to Gd-DTPA (when compared as R1=1/T1 values at 1 millimolar concentrations of Gd) and by multiples of 2.5 to 5.0 (when compared as the molarities of Gd required to produce a specified decrease in the T1 relative to a control solution (physiologic saline).
The reasons for using the latter method of comparison are that 1) millimolar concentrations of Gd are never achieved in vivo--actual tissue concentrations achieved in the usual image enhancement are between 20 and 100 micromolar Gd; 2) the slopes of R1 graphs are frequently nonparallel for different enhancing agents; 3) the second method allows agents to be compared according to the more customary means of chemical activity ratio, in other words, as the concentration required to produce a specified percentage decrease in the T1 (or T2) relaxation time. The second method is considered preferable. A drawback of conjugating DTPA to protein carriers for use in NMR image enhancement is that it has been difficult to stably conjugate more than 5 DTPAs (and hence Gd's) to each albumin molecule (Bulman et al. (1981) Health Physics V 40, p 228, Lauffer et al. (1985) Mag. Res. Imaging V 3, p 11 and Hnatowioh et al. (1982) Int. J. Appl. Radiat. Isot. V 33, p 327 (1982).
Comparably low substitution ratios (normalized for molecular weight) have been reported for immunoglobulins (Lauffer et al. (1985) Mag. Res. Imaging V 3, p 11 and Brady et al. (1983) Soc. Mag. Res., 2nd Ann. Mtg., Works in Progress, San Francisco, Calif.) and fibrinogen (Layne et al. (1982) J. Nucl. Med. V 23, p 627). This results from the relative difficulty of forming amide bonds, the comparatively low number of exposed amino groups on typical proteins which are available for coupling, and the relatively rapid hydrolysis of DTPA anhydride coupling substrate which occurs in the aqueous solvents required to minimize protein denaturation during conjugation (Hnatowich et al. (1982) Int. J. Appl. Radiat. Isot. V 33, p 327 (1982) and Krejcarek et al. (1977) Biochem. Biophys. Res. Comm. V 77, p 581). The overall effect of these suboptimal conditions is that a very large dose of carrier material is required to achieve significant in vivo effects on MR images. At this high dose, the carrier produces an unacceptable acute expansion of the recipient's blood volume by an osmotic mechanism. Indeed, low substitution ratios have generally limited the use of such protein-chelator-metal complexes to the more sensitive (low-dose), radiopharmaceutical applications (Layne et al. (1982) J. Nucl. Med. V 23, p 627).
An attempt to overcome this low substitution ratio has been made by conjugating DTPA to the non-protein carrier, cellulose (Bulman et al. (1981) Health Physics V 40, p 228), however the chemical method employed results in continued suboptimal substitution of DTPA to carrier, the nonbiodegradability of cellulose and its water-soluble derivatives and the reported molecular aggregation which results from organic-solvent conjugation (in dimethylformamide) of CNBr-activated cellulose to the diaminohexyl spacer groups which link the carrier to DTPA, have rendered this class of carrier-conjugates unacceptable for intravenous administration at the doses required for MR image enhancement.
A very important consideration in the image enhancement of solid tumors and inflammatory lesions by polymeric contrast agents is that, in order for these agents to extravasate (exit) efficiently from the microcirculation into adjacent diseased tissues, they must be completely soluble--e.g., not be contaminated by intermolecular or supramolecular microaggregates. Optimal tumor access and localization requires that the molecular size of such agents generally be less than approximately 2,000,000 daltons (ca. 2 to 3 nanometers in molecular diameter), and preferably less than 500,000 daltons (ca. 0.5 to 1.0 nanometers in molecular diameter) (Jain (1985) Biotechnology Progress V 1, p 8I). For this reason, with rare exceptions the particulate and microaggregate classes of contrast agents (which comprise the liposomes, colloids, emulsions, particles, microspheres and microaggregates, as described below) do not concentrate efficiently in most solid tumors or inflammatory lesions. Instead, following intravenous administration, these supramolecular-sized agents: a) are first circulated in the bloodstream for relatively short intervals (225 minutes to 24 hours, depending on size), potentially allowing direct image enhancement of the blood pool (plasma compartment); and b) are subsequently cleared by specialized (phagocytic) cells of the reticuloendothelial tissues (liver, spleen and bone marrow), potentially allowing selective enhancement of these normal tissues, but producing indirect (negative) enhancement of lesions within these tissues (due to exclusion of the agents from the diseased regions). Additionally, following installation into the gastrointestinal tract and other body cavities, these particulate and microaggregate classes of agents can produce direct image enhancement of the fluid within these cavities, and thereby potentially delineate mass lesions which encroach upon the lumens and cavities. Both microspheres and microaggregates are supramolecular in size. The microaggregate class of agents is produced (intentionally or unintentionally) by either a) molecular cross-linking of individual polymer molecules or b) secondary aggregation of previously single (soluble) polymers, as induce by charge attraction or hydrophobic bonding mechanisms. It is distinguished from the microsphere class of agents by virtue of its smaller particle size, which ranges from approximately 2,000,000 daltons (ca. 2 to nanometers in diameter) to 0.1 micrometers (=100 nanometers in diameter). It is important to note that microaggregates are cleared by reticuloendothelial phagocytes with significantly less efficiency and rapidity than are microspheres. In general, this property makes microaggregates a less preferred class of agents for visualizing the liver, spleen and bone marrow under the usual conditions of clinical imaging, for which prompt post-injection contrast enhancement is required.
(3) Gd-DTPA has been entrapped in liposomes (Buonocore et al. (1985) Proc. Soc. mag. Res. med. V 2, p 838) in order to selectively enhance images of the reticuloendothelial organs (liver, spleen and bone marrow) and potentially the lungs. Liver clearance is mediated by phagocytic (Kupffer) cells which spontaneously remove these small (0.05 to 0.1 um) particles from the bloodstream (Buonocore et al. (1985) Proc. Soc. Mag. Res. Med. V 2, p 838). (Particles larger than 3 to 5 um are selectively localized in the lungs due to embolic entrapment in lung capillaries.) A recent report indicates that the small-sized Gd-liposomes produce effective decrease in liever T1's (as determined spectroscopically without imaging) (Buonocore et al. (1985) Proc. Soc. Mag. Res. med. V 2, 0838). Also, insoluble GdDTPA colloids have recently been reported to enhance MR images of rabbit livers under in vivo conditions (Wolf et al. (1984) Radiographics V 4, p 66). However, three major problems appear to limit the diagnostic utility of these devices. The multilamellar, lipid envelopes of liposomes appear to impede the free diffusion of water protons into the central, hydrophobic cores of these carriers, as assessed by the higher does of Gd required for in vitro relaxivities equivalent to Gd-DTPA dimeqlumine (Buonocore et al. (1985) Proc. Soc. mag. Res. med. V 2, p 838). This increases the relative toxicity of each Gd atom.
Even more importantly, these same lipid components cause the carriers to interact with cell membranes of the target organs in a way which leads to a marked prolongation of tissue retention (with clearance times of up to several months) (Graybill et al. (1982) J. Infect. Dis. V 145, p 748 and Taylor et al., 1982 Am. Rev. Resp. Dis. V 125, p 610) . Two adverse consequences result. First, image enhancement does not return to baseline in a timely fashion. This precludes re-imaging at the short intervals (ca. 1 to 3 weeks) needed to assess acute disease progression and treatment effects. Second, significant quantities of the liposomally entrapped Gd-DTPA may be transferred directly into the membranes of host cells (Blank et al. (1980) Health Physics V 39, 913; Chan et al. (1985) Proc. Soc. Mag. Res. Med. V 2, p 846). This can markedly increase the cellular retention and toxicity of such liposomal agents. The consequences for Gd toxicity have not yet been reported. Protein (albumin) microspheres with entrapped Gd and Gd chelates have been prepared and determined (Saini et al. (1985) Proc. Soc. Mag. Res. Med. V 2, p 896) to have only modest effects on T1 relaxivity in vitro. This is because most of the Gd as well as other entrapment materials (Widder et al. (1980) Cancer Res. V 40, p 3512) are initially sequestered in the interior of these spheres and are released very slowly as the spheres become hydrated (with t1/2's of hours) (Widder et al. (1980) Cancer Res. V 40, 3512).
Emulsions of insoluble gadolinium oxide particles have been injected into experimental animals with significant image-enhancing effects on the liver (Burnett et al. (1985) Magnetic Res. Imaging V 3, p 65). However, these particles are considerably more toxic than any of the preceding materials and are thus inappropriate for human use. Because of the significant disadvantages of existing MR image contrast agents, the present applicant has formulated improved, second-generation prototype agents with reduced toxicity, phenominally increased effectiveness potentially increased selectivity of tumor and organ uptake, as well as a significant potential for enhancing blood flow images.
Many of the advantages shown for the present developments concerning NMR image-enhancing agents (also referred to herein as NMR contrast agents or MR (magnetic resonance) contrast agents) are also expandable to other areas. Gadolinium and related agents can produce characteristic changes in the NMR spectrum of adjacent NMR-susceptible nuclei. These changes include: modulation of resonance peak positions, widths, intensities, and relaxation rates (which affect intensity). Hence, perturbation of spectra by such chemical shift-relaxation agents can be used to localize and identify the source of NMR signal with respect to organ location, tissue compartment (intravascular versus extravascular), cell type within the tissue, and potentially, the specific metabolic pathways within cells which are altered by drugs and disease. Also in certain situations, ultrasound imaging or body scanning of radioisotopic emissions is particularly useful in achieving insight into internal structures. The radioisotopic emissions most frequency scanned are those of metallic radioisotopes emitting gamma particles, however, positron emission tomography is experiencing increased clinical use. The molecular formulation and mode of administering these radioisotopic metals will have significant consequences on the internal localization and body half-life of these radioisotopes, potentially leading to increased diagnostic usage of these ultrasound images and emission scannings.
The present invention includes an image-enhancing or spectral-shift agent comprising a biodegradable, water-soluble or insoluble melanin polymer, synthetic or derived from natural sources and having a signal-inducing metal or metal particle incorporated therein.
Melanins are a group of pigments derived from several different amino acid substrates in the natural world. Different types of melanins are responsible for much of the coloration in animals, plants, and bacteria. Melanins are usually classified in one of three groups: eumelanins, such as those found in hair, skin, feathers, and as a part of the coloration in reptiles and fish; phaeomelanins, which produce human red hair and the read fur of foxes; and allomelanins, which are most often present in bacteria and plants. An example of one of the most striking effects of a combination of eumelanins and phaeomelanins in the patterns found in the plumage of certain tropical birds.
Melanins are created by the action of enzymes on any of several amino acid precursors. For example, tyrosine is the precursor of most eumelanins. However, the exact process by which an amino acid is converted to melanin is unknown, and indeed, seems to vary from one pigment to another even when formed from the same substrate. In general, black pigments are formed from precursors including 3,4-dihydroxyphenylalanine (DOPA), catechol, and various other dihydroxy- substances.
______________________________________ OH CH COOH OH CH OH NH OH NH 3,4-dihydroxyphenylalanine 5,6-dihydroxyindole (DOPA) HO OH OH OH catechol 1,8-dihydroxynaphthalene ______________________________________
These substances have many active centers for polymerization. Compounds with fewer active centers are precursors to melanins of brown, reddish-brown, or yellow-brown color.
Melanins in pure form are usually insoluble in water as well as in most organic solvents, making them difficult to work with. In part due to this, not all of the chemical properties of melanins have been identified. However, recent studies using electron spin spectroscopy have identified free-radical properties of melanin.
Synthetic melanins can be created from most of the precursors used in nature by using almost any chemical oxidant or free-radical polymerizer, or, in some cases, merely leaving dissolved substrate open to the air overnight.